Scintillator compositions used in the detection of ionizing radiation respond to the reception of an energetic photon by generating a pulse of scintillation light. The energetic photon may for example be an x-ray or gamma photon, and the resulting light pulse typically comprises a plurality of optical photons having wavelengths in the infrared to ultraviolet spectral range. The scintillation light pulse is conventionally detected using an optical detector which generates an electrical pulse at its output that may subsequently be processed by electronic circuitry. Together, a scintillator composition that is optically coupled to an optical detector, referred to herein as an ionizing radiation detector, may be configured to receive energetic photons from the imaging region of for example a PET or a SPECT imaging system. In such an imaging system the received energetic photons provide data for subsequent reconstruction into an image indicative of radioactive decay events in the imaging region. Ionizing radiation detectors may likewise be used to detect energetic photons traversing the imaging region of a CT imaging system.
The selection of a scintillator composition for use in such ionizing radiation detectors is subject to a variety of constraints imposed by the imaging system in which the detector is to be used. A PET imaging system may be considered to impose some of the most stringent limitations on such scintillator compositions owing to the need to accurately time the reception of each individual gamma photon, and the need to determine its energy. The time of reception of each gamma photon is important in the identification of coincident pairs of gamma photons, which by virtue of their detection within a narrow time interval are interpreted as having been generated by a radioactive decay event that lies along a line in space between their detectors, termed a line of response. The energy of each gamma photon may be further used to confirm whether timewise-coincident pairs of gamma photons share a common origin along the line of response by discarding pairs of events in which one of the gamma photons has undergone energy-altering scattering. Coincident pairs of gamma photons each having an energy that is within a predetermined narrow range are interpreted as being the product of a single radioactive decay event. The energy of each gamma photon is typically determined by integrating the light pulse, or by counting the number of scintillation photons produced by each gamma photon.
These requirements manifest themselves in the need for a scintillator composition with a high light yield and a short decay time. The high light yield improves the signal to noise ratio at the output of the optical detector, and the short decay time is important in the prevention of pile-up. Pile-up occurs when the scintillation light pulse from a gamma photon overlaps in time with the scintillation light pulse generated by a previously-received gamma photon. Pile-up degrades both the timing accuracy of a gamma photon detector and the ability to determine the energy of each gamma photon.
The light yield of a scintillator composition, measured in units of photons per MeV, is a measure of the number of scintillation photons produced in response to a received energetic photon. A high light yield is achieved by using a scintillator composition with strong photopeak absorption. The decay time of a scintillator composition is conventionally determined by fitting the decay of the scintillation light pulse with two time constants. A primary decay constant accompanied by the percentage of the total scintillation light emitted during the primary decay period extrapolated to infinity models the initial decay, and this is followed by a secondary decay constant that is likewise accompanied by the percentage of the total scintillation light emitted during the secondary decay period extrapolated to infinity. Thus, the light yield, and the decay time, comprising its primary and secondary components, are two parameters that affect the sensitivity of a scintillator composition.
A PET imaging system typically comprises an imaging region, around which are disposed a plurality of ionizing radiation detectors, or more specifically, gamma radiation detectors, which in combination with timing circuitry are configured to time the reception of gamma photons. The gamma photons may be produced following the decay of a radiotracer within the imaging region. Pairs of gamma photons that are detected within a time interval of typically +/−3 ns of each other are deemed coincident and are interpreted as having been generated along a line of response (LOR) between their two gamma radiation detectors. Multiple lines of response are subsequently used as the data input to a reconstruction processor executing reconstruction algorithms to reconstruct an image indicative of the radiotracer distribution within the imaging region.
The timing certainty associated with the detection of a pair of gamma photons in a PET imaging system is determined by its coincidence resolving time, or CRT. The CRT is the narrowest time interval within which two gamma photons received simultaneously by different gamma photon detectors are certain to have been detected. A coincidence resolving time of less than +/−3 ns is typically desirable in a PET imaging system in order to be assign their detection to a 90 cm LOR. This corresponds to approximately the maximum typical bore diameter of a PET imaging system. In a PET imaging system having a CRT of shorter than +/−3 ns, an estimate of the originating position of a pair of gamma photons along the line of response may further be determined based on the exact times of detection of two gamma photons. A shorter CRT permits the reconstruction processor in such an imaging system to reconstruct a more accurate image of the radiotracer distribution within its imaging region. This principle is employed in a Time of Flight (TOF) PET imaging system, in which a CRT of less than +/−3 ns is preferred, and in which a CRT of less than +/−1 ns is even more preferred. The CRT of a PET imaging system is affected by the scintillator decay time, by the probability of gamma photon scattering within the scintillator composition, by the depth of interaction of a gamma photon within the scintillator composition, by the geometry of the scintillator composition, and by the timing accuracy of the electronic timing circuitry. Minimizing the CRT is therefore a key goal in improving the sensitivity of a PET imaging system and this places additional emphasis on using a scintillator composition with a short decay time.
Some known scintillator compositions that are suitable for use in PET imaging systems are disclosed in publication “Luminescence: From Theory to Applications”, Wiley-VCH, Darmstadt, 2007, C. Ronda (Ed.). These include LYSO (lutetium yttrium oxy-orthosilicate) crystals and LaBr3 crystals. Scintillation in LYSO has been reported with a light yield of 33000 photons/ MeV in the presence of a decay time of 44 ns. A higher light yield has been reported in LaBr3 in the presence of a decay time of 25 ns. Single crystal compositions are investigated almost exclusively owing to the best combination of stopping power, decay time and light yield being found in single crystal compositions.
More recently, garnet crystals have been reported for use as scintillator compositions as disclosed in publication “Composition Engineering in Cerium-Doped (Lu,Gd)3(Ga,Al)5O12 Single-Crystal Scintillators”, K. Kamada et al, Cryst. Growth Des., 2011, 11 (10), pp 4484-4490. According to this publication, the adapted results of which are illustrated in the table in FIG. 1, the light yield of a reference single crystal Gd3Ga2Al3O12 garnet composition is significantly reduced from 45931 photons/MeV to 30627 photons/MeV when lutetium is incorporated into the composition to form Lu1Gd2Ga2Al3O12. The light yield in the table in FIG. 1 can also be seen to reduce from 30627 photons/MeV to 18166 photons/MeV when the relative gallium-to-aluminium content is increased. This effect is known in the literature as thermal quenching.
Garnet scintillator compositions are also known from documents US6630077B2, DE102013103783A1, US7252789B2, US6793848B2 and EP463369A1.
The present invention seeks to overcome limitations of known scintillator materials.